Magnetic resonance imaging (MRI) detects the faint nuclear magnetic resonance (NMR) signals given off by protons in the presence of a strong magnetic field after excitation by a radio frequency signal. The NMR signals are detected using antennae termed “coils”. The term “coil” is also commonly used to refer to the antenna(e) and its housing or support structure. Thus “coil” may refer to a structure that contains a number of coils. “Coil element(s)” is used to refer to the electrical part of the device, the radio frequency coil or antennae.
NMR signals are extremely faint. Sensitivity of a coil to these signals decreases rapidly with increasing distance between the coil and the volume of interest. Coils such as FIG. 8, or butterfly coils, solenoid coils, volume coils, and surface coils are therefore placed in close proximity to the region of interest of the imaged object. The size of the local coils is kept small to allow them to be easily fit to the patient on the MRI patient table and to enable imaging of only the imaging volume of interest, since imaging regions that are not required adds noise to the acquired signal unnecessarily. Coils local to the anatomy of interest tend to have a higher signal-to-noise ratio (SNR) than larger coils such as a “body coil” which is useful for obtaining large survey scans of the patient.
The smaller the size of the local coil, the smaller its field of view, or sensitivity profile. Imaging of larger areas using the smaller coils requires the use of multiple small coils, either simultaneously in a combined manner or by moving the coil between imaging acquisitions.
Coils can be operated individually, as multiple coils in a phased array, circularly polarized or in quadrature mode. Combining signals from multiple coils can yield improvements in SNR. Part of the challenge associated with using multiple coils for imaging is the fact that the fields of individual coils may interact, resulting in coil-to-coil coupling, where these interactions serve to reduce the coil quality factor, or Q. In the prior art various patents have been presented whose proposed objective is to reduce this coupling. One technique for reducing the coil-to-coil coupling of a multi-coil array is to overlap adjacent coils by approximately 10% of their area, such that their additional field contributions cancel resulting in no coupling. In cases where there may be more than 2 coils, the process of decoupling by overlapping can be complicated, as coil coupling may occur between non-nearest neighbors in which the field cancellations are complicated significantly. In these cases, coupling can be reduced through the addition of capacitors, inductors or additional circuits between coils which experience some amount of coupling. Low-impedance preamplifiers may be added to the coil system which can reduce the effects of coil coupling. In much of the prior art, combinations of these various techniques have been described and employed successfully.
Furthermore, operating a coil as a receive-only coil requires that the coil is blocked, or uncoupled from the magnet body coil while the body coil or other coils are acting in transmit, or excitation mode. Again, various patents are presented in the prior art that seek to improve upon this process.
A further consideration with coil systems is their ability to operate in a parallel imaging mode. In these modes of operation, imaging techniques such as SMASH, SENSE, PILS or GRAPPA, require coils to be imaging independent volumes. Based on the sensitivity profiles of these coils operating independently, a reconstruction algorithm can be implemented that enables reconstruction of a full image volume in a fraction of the conventional image acquisition time. Coils should image independent volumes for optimal parallel imaging, and therefore decoupling strategies that employ overlapping of coils are non optimal.
An additional consideration with coil technologies and generally with MRI systems is the push towards higher and higher numbers of simultaneous imaging channels. Coil systems are routinely implemented in 8-channel systems utilizing eight separate antennas, but some systems currently implement up to 96 channels, with higher numbers of channels being planned. The benefits of additional channels include higher acceleration factors from parallel imaging implementations, and smaller coils for higher signal to noise ratios. With this constant drive to upgrade MRI systems, legacy coil systems become obsolete. There is currently no way of upgrading the number of channels associated with a coil without buying an entire coil apparatus which includes the main structure, the coil circuitry and the connection to the MRI.
Coils must be tuned to the Larmor frequency associated with the magnet field strength in which it is meant to operate, i.e. 1.5 T requires a coil tuned to 63.86 MHz, 3.OT requires 127.7 MHz. In most commercial coils, the coil elements or antenna are inseparable from the patient support structure, or are inseparable from the coil housing. For each MRI having a different field strength, therefore, a new coil system, consisting of the coil elements, coil housing, patient support (if any) and cabling is required.
Although most MRI imaging is concerned with signals from hydrogen atoms, other nuclei (e.g. CI3K, P, Na) are sometimes of interest for MR spectroscopy or imaging. Traditionally, the low signal to noise ratio associated with measurement of these nuclei have precluded their use for any practical clinical imaging. However, with the advent of improved coil technologies and higher field strengths, these techniques are becoming more practical. However, coils tuned to the appropriate precessional (Laramor) frequencies of the nuclei are required as well as the associated circuitry to enable acquisition of these signals on the limited bandwidth of standard MRI systems. There are some systems that are appropriate for double tuned imaging (i.e. hydrogen in combination with another species), however there are limitations associated with integrating these multiple coils into one coil housing.
Furthermore, coil switching, multiplexing, or dynamic coil selection strategies enable activation and inactivation of sets of coils from a larger coil array set. This strategy can be used to optimize a subset of coils for imaging of anatomies of a smaller volume, or to switch between areas of interest during the image acquisition or imaging procedure. For this strategy to be accommodated, the set of coils, or subset of these coils must be appropriately designed. There currently are no systems which accommodate this type of imaging strategy with a modular coil system design.
Another consideration for imaging, particularly of the human breast, is the varying positions in which the breast is imaged for MRI, US and mammography and in which surgical interventions may be performed. As the breast may often be imaged in the prone position for MRI while surgery and ultrasound (US) imaging examinations are performed almost exclusively in the supine position, it is difficult to correlate anatomical features between these positions. MR imaging of the breast in the supine position is very difficult and has not been accomplished with any degree of success.
As coil technology complexity and clinical demands continue to increase, a new strategy for coil systems is needed. A system is required that permits sizing coils appropriate to the anatomy, maximizes the number of coils used, providing coils tuned for different field strengths and nuclei, optimizes parallel imaging configurations, as well as providing an upgrade path for accommodating greater numbers of imaging channels.